Field stimulation about a discontinuity of the myocardium to capture the heart at reduced pacing thresholds

ABSTRACT

Improved pacing thresholds for capturing the heart are achieved by forming a discontinuity in the cardiac tissue of the heart chamber, disposing a pacing electrode at a distance exceeding a space constant of the cardiac tissue from the discontinuity in the cardiac tissue, and applying a stimulus of a first polarity at an energy insufficient to cause the directly stimulated tissue adjacent to the pacing electrode to propagate a depolarization wave through the cardiac tissue mass of the heart chamber but sufficient to induce a transmembrane potential change at the tissue adjacent to the discontinuity that results in a propagated wave front. Thus, pacing energy is advantageously reduced.

RELATED APPLICATION

This application claims priority and other benefits from U.S.Provisional Patent Application Ser. No. 60/337,273 filed Dec. 3, 2001,entitled PACING VIA VIRTUAL ELECTRODES FORMED AROUND A FIELD STIMULATEDMICRO-LESION

FIELD OF THE INVENTION

The present invention relates to methods and electrode configurationsfor pacing the heart, particularly by pacing the heart at reduced pacingenergy through inducement of virtual anodes and cathodes along lesionsformed in the heart tissue.

BACKGROUND OF THE INVENTION

Many implantable medical devices (IMDs) have been developed over theyears for clinical implantation in patient's bodies that deliverelectrical stimulation to a body organ, muscle, nerve or brain cells.Each year approximately 750,000 patients develop bradycardia symptomssuch as dizziness, extreme fatigue, shortness of breath, or faintingspells. These symptoms are caused by abnormally slow or irregular heartrate, and the most effective method to relieve these symptoms is toimplant a pacing system that generates and delivers pacing pulses to asite in or adjacent to a heart chamber. Pacing systems are incorporatedinto a wide variety of implantable pacemakers and also into implantablecardioverter defibrillators (ICDs). Such pacing systems comprise animplantable pulse generator (IPG) and one or more lead interconnectingthe IPG circuitry with pace/sense electrodes implanted against or intothe myocardium of the heart.

Each heart cell contains positive and negative charges due to theselective permeation of certain ions, such as potassium and sodiumthrough the cell membrane. When the cell is at rest, the inside of thecell is negatively charged with respect to the outside. The negativecharge is dissipated when the cell is disturbed by an electrical signalthat causes the permeability of the cell membrane to change and allowsthe ingress of positive charge ions. The resulting dissipation of thenegative charges constitutes the “depolarization” of the cell.Simultaneously, the cell contracts causing (in conjunction with thecontraction of adjoining cells) the heart muscle to contract. Thus, thestimulation of the heart muscle affects both the depolarization and thecontraction of the once-polarized myocardial cells that make up themuscle.

Following depolarization and contraction of a heart cell, the“repolarization” or recovery of the cell commences so that the cell isready to respond to the next applied stimulus. During the repolarizationtime interval, the cell membrane begins to pump out the positive-chargedions that have entered following the application of the stimulus, thatis, during the depolarization of the cell. As these positive chargesleave, the inside of the cell membrane starts to become negative again,the cell relaxes, and the potential difference builds up again.

The individual myocardial cells are arranged to form muscle fibers andsheets that, in gross, constitute the heart itself. The depolarizationof the atrium is characterized by a P-wave viewed on anelectrocardiogram (ECG), and depolarization and repolarization signalsof the ventricle, are referred to as the QRS complex and the T wave,respectively. The sequence of depolarization, which manifests itself ina contraction of the heart muscle, and repolarization, which manifestsitself in the relaxation and filling of the interior chambers of theheart with blood, is accomplished through a system of specialized muscletissue that functions like a nerve network. Depolarization signals aregenerated in the SA node of specialized cardiac cells located in theatria at a rate that is appropriate for the body's physiologic demandfor cardiac output. The system then conducts these impulses rapidly toall the muscle fibers of the ventricles, ensuring coordinated,synchronized pumping.

When this system fails, or is overridden by abnormal mechanisms, apacing system may be needed to generate and deliver trains of pacingpulses through pace/sense electrodes to maintain proper heart rate andsynchronization of the filling and contraction of the atrial andventricular chambers of the heart. The pacing circuitry of pacemaker andICD IPGs is powered by a battery, and each delivered pacing pulseconsumes a discrete bolus of the battery energy. Consequently, the IPGlongevity is primarily governed by the battery lifetime. Currently, theIPG longevity can range from approximately 3 to 10 years depending onthe type of IPG (e.g., pacemaker or ICD). The IPG must be replaced whenthe battery is depleted, an expensive procedure that also posessignificant discomfort and risk to the patient.

The pacing current drawn by each pacing pulse is a major factor thatimpacts the battery life and device longevity, although its impact isgreater for some devices than the others. For example, recentlydeveloped bi-ventricular pacing systems incorporated into pacemakers andICDs present a high current drain since two pacing pulses must bedelivered to synchronously pace both ventricles at a pacing rate thattypically depends upon the patient's physiologic need for cardiac outputas determined by an activity sensor, for example. Thus, the reduction indelivered pacing current would certainly increase the IPG longevity andcould allow the battery and corresponding IPG size to be reduced, andtherefore positively impact lives of thousands of patients receivingbattery powered IMDs.

In the history of implantable cardiac pacemakers, great strides havebeen made in increasing longevity, reliability, and versatility of IPGsand the associated lead systems. In the early days of implantablecardiac pacemakers, battery depletion was rapid, leading to exhaustionof the IPG batteries within a year from implantation. The high energyconsumption was due to a wide variety of factors, including battery selfdischarge, pace/sense electrode-tissue interface inefficienciesrequiring delivery of high energy pacing pulses, and high currentconsumption by discrete electronic circuit components.

It was recognized from the outset of cardiac pacing that IPG batterycurrent drain is directly proportional to the amount of energy that isnecessary when delivered to the heart to cause the heart to depolarize,i.e., to “capture” the heart. Over the last forty years, reliability andlongevity have dramatically improved due to improvements in batterytechnologies, lead and pace/sense electrode technologies, electroniccircuitry current consumption and a wide variety of other areas. Asimprovements in one area led to increased longevity and reliability,attention was focused on the other areas.

In this evolutionary process, early studies were conducted to determineif the optimum stimulation pulse polarity and wave shape could be foundthat would achieve capture of the heart at the lowest expenditure ofpulse energy in order to prolong pacemaker battery life as reported, forexample, by Egbert Dekker, M.D., in “Direct Current Make and BreakThresholds for Pacemaker Leads”, (Circulation Research, vol. XXVII,November 1970, pp. 811–823). In the infancy of cardiac pacemakers,experiments were performed using various forms of electrical stimulationpulses including anodal (positive going) and cathodal (negative going)pacing pulses having pulse energy exceeding the stimulation threshold totrigger depolarization of myocardial cells.

Contemporaneously, attention was focused on other factors, particularlypace/sense electrode technologies, high energy density, low selfdischarge, battery technology, variable pulse energy output pulsecircuits, and capture threshold determination techniques that madedramatic improvements in IPG longevity, reliability and size. Thepace/sense electrode technologies have included pace/sense electrodematerials including substrates, coatings and surface treatments,pace/sense electrode shapes, pace/sense electrode surface areas, andpace/sense electrode configurations as well as minimizing local tissueinjury when the pace/sense electrode fixed in place by a tissuepenetrating active fixation mechanisms, delivery of steroids to thestimulation site by incorporation of steroid eluting elements in thelead body adjacent to the fixation mechanism or coatings on the fixationmechanism.

Today's implantable pacemakers and pacing systems incorporated into ICDsare far more versatile and offer a wider variety of therapies formedical conditions that were not imagined in the infancy of cardiacpacing. Currently, electrical stimulation generated by a pacemaker orICD IPG is in the form of pacing pulses typically having a fixedduration in the order of about 0.5 ms, a voltage of less than 5 volts,and a resulting delivered current dependent upon the collectiveimpedance or load that the pulse is delivered through a cardiac leadconductor and the pace/sense electrode-tissue interface at an activepace/sense electrode. The exponential decaying voltage, cathodal(negative going) pacing pulse shape achieved by a relatively simple,monophasic capacitive discharge output circuit has become accepted asthe standard pacing pulse for many years.

Thus, a negative voltage pulse is typically delivered at the activepace/sense electrode, whereby the active pace/sense electrode ischaracterized as a cathode pace/sense electrode and the return orindifferent pace/sense electrode in the discharge path is characterizedas an anode pace/sense electrode. A cathodic electrical field ofsufficient strength and current density has to be impressed upon theexcitable tissue in the vicinity of the active site to initiateconduction of a depolarization wave through the entire cardiac tissuemass of a heart chamber that causes the heart chamber to contract andexpel blood from the heart chamber, i.e., to capture the heart. Theminimum pacing pulse energy necessary to produce that effect is referredto as the “stimulation threshold” or “pacing threshold.” The greater theefficiency of the cathode in impressing the electric field on thetissue, the smaller is the amplitude and/or duration of the pulserequired to exceed the stimulation threshold. With the widespreadadoption of multi-programmable parameters including programmable pulsewidth and amplitude, physicians have become accustomed to determiningthe patient's pacing threshold and setting the energy level to a minimumvalue to capture the heart plus an adequate safety margin.

Despite these efforts and realized improvements, a need remains tofurther reduce pacing thresholds.

SUMMARY OF THE INVENTION

The present invention provides improved pacing thresholds throughadoption of technologies that heretofore have not been employed in theprovision of pacing via pace/sense electrodes to a heart chamber.

Improved pacing thresholds for capturing the heart are achieved byforming a discontinuity in the cardiac tissue of the heart chamber,disposing a pacing electrode at a distance exceeding a space constant ofthe cardiac tissue from the discontinuity in the cardiac tissue, andapplying a stimulus of a first polarity at an energy insufficient tocause the directly stimulated tissue adjacent to the pacing electrode topropagate a depolarization wave through the cardiac tissue mass of theheart chamber but sufficient to induce a transmembrane potential changeat the tissue adjacent to the discontinuity that results in a propagatedwave front. Thus, pacing energy is advantageously reduced employing themethods and apparatus of the present invention.

The present invention is preferably implemented in unipolar and bipolarpacing leads having an electrode head at the lead body distal endwherein the electrode head supports at least one active, cathodal pacingelectrode to bear against or be disposed into the myocardium at thespace constant distance from the discontinuity. In a bipolar pacinglead, the electrode head supports the active, cathodal pacing electrodeand the indifferent, anodal pacing electrode with the discontinuityformed between the pacing electrodes at the space constant distance fromeach pacing electrode.

The lesion or cleft is preferably created by a non-conductive cuttingblade or fixation screw supported by the electrode head to be directedinto the myocardium upon fixation of the electrode head against theendocardium or epicardium.

This summary of the invention and the advantages and features thereofhave been presented here simply to point out some of the ways that theinvention overcomes difficulties presented in the prior art and todistinguish the invention from the prior art and is not intended tooperate in any manner as a limitation on the interpretation of claimsthat are presented initially in the patent application and that areultimately granted.

BRIEF DESCRIPTION OF THE DRAWINGS

Other advantages and features of the present invention will be readilyappreciated as the same becomes better understood by reference to thefollowing detailed description when considered in connection with theaccompanying drawings, in which like reference numerals designate likeparts throughout the figures thereof and wherein:

FIGS. 1A and 1B are graphical depictions of anodal and cathodal,intracellular stimulation;

FIGS. 1C and 1D are graphical depictions of anodal and cathodal,extracellular stimulation;

FIGS. 2A and 2B are schematic illustrations of excitation of anisotropic cardiac tissue;

FIGS. 3A and 3B are schematic illustrations of excitation of realisticanisotropic cardiac tissue with unequal anisotropy ratio in theextracellular and intracellular domain;

FIGS. 4A and 4B are graphical depictions of steady state polarization ofa cardiac fiber;

FIGS. 5A–5D are graphical depictions of the formation of virtual sourcesbracketing an intracellular discontinuity;

FIGS. 6A and 6B are schematic illustrations of excitation of anisotropiccardiac tissue inducing virtual sources bracketing an intracellulardiscontinuity;

FIG. 7 is a schematic illustration of an experimental setup fordetermining pacing thresholds in cardiac tissue prior to and followingforming a lesion in the cardiac tissue;

FIGS. 8A–8C are schematic illustrations depicting stimulation sites ofheart chambers of hearts stimulated using the test setup of FIG. 7;

FIGS. 9A–9C are tracings of the cardiac ECG as well as the appliedpacing pulses and depolarization responses obtained from a heartstimulated using the test setup of FIG. 7;

FIGS. 10A and 10B are graphical depictions of pacing threshold dataobtained using the test setup of FIG. 7 for unipolar stimulation appliedprior to and following formation of linear lesions in guinea pig hearts;

FIGS. 11A and 11B are graphical depictions of pacing threshold dataobtained using the test setup of FIG. 7 for bipolar stimulation appliedprior to and following formation of linear lesions in guinea pig hearts;

FIGS. 12A and 12B are graphical depictions of time dependent pacingthreshold data obtained using the test setup of FIG. 7 for unipolar andbipolar stimulation prior to and following formation of linear lesionsin guinea pig hearts;

FIGS. 13A and 13B are graphical depictions of the percent reduction inunipolar and bipolar pacing thresholds following formation of lesions inguinea pig hearts subjected to threshold testing using the test setup ofFIG. 7;

FIGS. 14A and 14B are graphical depictions comparing unipolar andbipolar pacing thresholds prior to and following formation of lesions inguinea pig hearts subjected to threshold testing using the test setup ofFIG. 7;

FIG. 15 is a plan view of a first embodiment of a pacing leadincorporating a retractable and extendable cutting blade made from aninsulating material for forming a discontinuity in cardiac tissuebetween anodic and cathodic pacing electrodes disposed on an electrodehead distal end;

FIG. 16 is an expanded end view of the electrode head distal end of thelead of FIG. 15;

FIG. 17 is an expanded side view of the electrode head distal end of thelead of FIG. 15;

FIG. 18 is an expanded side view in partial cross-section of theelectrode head of the lead of FIG. 15 depicting the cutting bladeextended to form a discontinuity in cardiac tissue between the anodicand cathodic pacing electrodes;

FIG. 19 is an expanded side view in partial cross-section of theelectrode head of the lead of FIG. 15 depicting the cutting bladeretracted into a chamber of the electrode head during transvenousadvancement to an implantation site;

FIG. 20 is a side view of an electrode head of an epicardial pacing leadthat supports a cutting blade to form a discontinuity in cardiac tissuebetween the anodic and cathodic pacing electrodes;

FIG. 21 is a bottom view of the electrode head of FIG. 20;

FIG. 22 is a side view of an electrode head of an epicardial pacing leadthat supports a solid screw to form a discontinuity in cardiac tissuebetween the anodic and cathodic pacing electrodes; and

FIG. 23 is a bottom view of the electrode head of FIG. 21.

The drawing figures are not necessarily to scale.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS OF THE INVENTION

In the following detailed description, references are made to exemplaryembodiments for carrying out the invention in reducing pacing thresholdsof chronically implanted endocardial or epicardial pacing leads. It isunderstood that other embodiments may be utilized without departing fromthe scope of the invention. The invention and its preferred embodimentmay be implemented in unipolar, bipolar or multi-polar, cardiac pacingleads having one or more pace/sense electrode(s) at or adjacent the leadbody distal end. The lead is adapted to be coupled to the connectorassembly of an implantable pulse generator (IPG) for pacing the heartthough the pace/sense electrode(s).

In seeking to lower pacing thresholds, it is first advisable to explorehow applied pacing pulses cause the heart to depolarize the heart.Below, we first discuss the electrical excitation of a single cell, thebasic unit of cardiac tissue, and then discuss the response to appliedstimulation at the tissue level. For simplicity, we start with adiscussion of pacing by treating the myocardium as an isotropic medium.This is followed by discussion of pacing in the context of morerealistic (but more complex) anisotropic bidomain model. See, Roth etal., “A bidomain model for the extracellular potential and magneticfield of cardiac tissue”. IEEE Trans Biomed Eng. 1986;33:467–9. Then, wediscuss the concept of virtual sources and virtual electrodes that wehave explored experimentally to develop new pacing methods andelectrodes of the present invention to lower the pacing threshold by anappreciable amount, e.g., 50% or more, of the currently attained values.

Excitation of a Single Cell

A typical mammalian cardiac cell is cylindrical in shape, approximately120 μm in length and 20 μm in diameter. A single cardiac cell can beexcited either by intracellular stimulation illustrated in FIGS. 1A and1B or by extracellular stimulation illustrated in FIGS. 1C and 1D. InFIGS. 1A–1D, the horizontal axis is space or distance, and the verticalaxis is either extracellular potential (φ_(l)) or intracellularpotential (φ_(e)) or transmembrane potential (V_(m)). Note that anelectrode of a given polarity can depolarize or hyperpolarize the celldepending on its extracellular or intracellular location.

In FIGS. 1A and 1B, a micro-electrode (e.g. a micropipette) is used toimpale the cell membrane so that the micro-electrode penetrates into andis continuous with the intracellular space. The intracellularmicro-electrode can then be used to inject or withdraw current from thecell. The intracellular micro-electrode can be referred to as an “anodalelectrode” when positive (anodal) current is injected into the cardiaccell and as a “cathodal electrode” when negative (cathodal) current isinjected into the cardiac cell.

For example, in FIG. 1A, anodal current is injected into the cellthrough the intracellular micro-electrode that raises the intracellularpotential φ_(l) relative to the extracellular potential φ_(e). Theincrease in intracellular potential is uniform along the cell lengthbecause the cell is small in size and has high intracellularconductivity. As a result, the transmembrane potential (V_(m)=φ₁−φ_(e))is raised, and the cell is depolarized. The conductance of sodiumchannels (small proteins spanning the cell membranes) undergoes a veryrapid increase when the transmembrane potential V_(m) attains athreshold value (˜−60 mV from resting value of ˜−90 mV). Sodium ionsthen flood the intracellular space and further raise the transmembranepotential V_(m) resulting in a cascade of time-dependent andvoltage-dependent changes in the conductance of other cell membranechannels (e.g., Ca²⁺ and K⁺ channels) causing the cell to fire an actionpotential.

However, the cell is hyperpolarized and not depolarized if cathodalcurrent is injected into the cell through the intracellularmicro-electrode as shown in FIG. 1B. The intracellular potential φ_(l)is lowered relative to the extracellular potential φ_(e). Again, thedecrease in intracellular potential is uniform along the cell lengthbecause the cell is small in size and has high intracellularconductivity.

As shown in FIGS. 1C and 1D, a single cardiac cell can be stimulatedwith an extracellular electrode disposed outside the cell membrane. Theapplication of positive (anodal) stimulation through the extracellularelectrode results in a positive extracellular potential φ_(e) that fallsmonotonically with distance from the electrode. The intracellularpotential φ_(l) is raised to a uniform value that is a weighted averageof extracellular potential φ_(e) around the cell. Thus, thetransmembrane potential V_(m) has the profile shown in FIG. 1C. Notethat the center of the cell now is hyperpolarized. Thus, the excitationwould occur only if the depolarization at the ends of the cell exceedsthe threshold value, and would require a large extracellular current.If, however, the polarity of the electrode is reversed to apply negative(cathodal) stimulation, the transmembrane potential V_(m) profile isreversed as shown in FIG. 1D. In this case, the cell center isdepolarized resulting in easier cell excitation.

Thus, a cardiac cell can be excited by anodal or cathodal stimulationapplied through respective anodal or a cathodal stimulating electrodes,depending on stimulating electrode location (extracellular versusintracellular).

Modeling Myocardium as an Isotropic Medium

To date, chronically implanted pace/sense electrodes (i.e., electrodesemployed to deliver pacing pulses and to sense intrinsic heart signals)coupled with pacemaker or ICD IPGs are not designed to and cannotpenetrate a viable cell membrane to apply intracellular stimulation to asingle cardiac cell. Any penetration of cardiac cells that may occurduring implantation or fixation of a pace/sense electrode into themyocardium irreparably damages the cell(s) and results in scar tissuecontacting the electrode or fixation mechanism. Therefore, whenpacemaker or ICD IPGs generate and deliver narrow pulse width (˜0.5 ms),negative going or cathodal pulses to a pacing site of the heart in orderto pace a heart chamber, the pacing current is injected in theextracellular space. The cathodal pacing pulse is generated by dischargeof a capacitor, typically charged to a voltage of 5 volts or lessbetween discharges, through a discharge circuit or load. The dischargeload comprises the lead conductor(s), the cathodal, active pace/senseelectrode at the pacing site, an anodal return or indifferent pace/senseelectrode, the cardiac and other body tissues and fluids between theactive and indifferent pace/sense electrodes, and the electrode-tissueinterfaces at the electrode surfaces. The active pace/sense electrode islocated typically at the distal end of a cardiac lead, and theindifferent pace/sense electrode is either located on the same lead orlocated more remotely, typically on the conductive housing of the IPG.The active pace/sense electrode either is fixed to bear against theendocardial or epicardial heart surface (referred to as passivefixation) or penetrates through the endocardial or epicardial heartsurface into the myocardium (referred to as active fixation).

For a single cardiac cell that is stimulated with an extracellularelectrode shown in FIG. 1D, the pacing pulse must raise thetransmembrane potential, Vm, of a critical length of the cell above athreshold value to cause a regenerative action potential. However, inpractice, for a 3-dimensional cardiac tissue the cathodal stimulationenergy applied to the extracellular domain through the typical activepace/sense electrode must depolarize a critical volume of the tissue asillustrated in FIGS. 2A and 2B so that the depolarized volume of cellsacts as a foci of depolarization of the entire heart chamber. A pacingpulse having a pulse energy exceeding the threshold value and causingthe heart chamber to depolarize is said to “capture” the heart. Thereason that a depolarization wave produced by a pacing pulse ofthreshold energy is able to invade the entire heart is that the cardiactissue is an electrical syncytium in which every cardiac cell isconnected to the next cardiac cell via intercellular gap junctions(i.e., small pore-like proteins structures that connect two adjacentcells) as shown in the inset to FIG. 2A. Thus, a trans-membrane currententering into one cardiac cell can be dissipated into an adjacentcardiac cell by electrotonic interaction (source-sink interaction). Ifthe volume of cardiac tissue that is caused to depolarize by the appliedcathodal stimulus is below a threshold volume, the current sink is largerelative to the source; therefore, the applied excitation fails toresult in a conducted depolarization through the cardiac muscle as shownin FIG. 2B. The current sink from the adjacent tissue is so large thatthe tissue excitation is suppressed. The source-to-sink mismatchdecreases with an increase in the amount of tissue directly depolarizedby the applied cathodal stimulus. Thus, the applied pacing pulse energymust provide enough current to excite a critical mass of cardiac tissueto result a conducted depolarization to occur that captures the heart.

The amount of cathodal stimulation energy required to excite thecritical mass of cardiac cells illustrated in FIG. 2A that in turncaptures the heart is referred to as the stimulation threshold or pacingthreshold. In practice, the pacing pulse energy (pulse width or pulsevoltage) is periodically adjusted by an auto-threshold algorithm or byprogramming so that the applied pacing pulse energy exceeds the pacingthreshold by a sufficient safety margin to conserve battery energy.

From the above discussion any strategy that can decrease source-sinkinteraction would decrease the critical mass required for excitation,and consequently will decrease current consumption and prolong batterylife. In accordance with the present invention, we manipulate thecardiac tissue precisely to accomplish this goal. But before we examinehow such manipulation of cardiac tissue can be accomplished, we mustintroduce a slightly more complex but more accurate anisotropic model ofthe cardiac tissue.

Modeling Myocardium as an Anisotropic Medium

The heart as a whole is much more complex than the cells or cell massesdepicted in FIGS. 1A–1D and 2A–2B. The cardiac cells that contract andrelax in the normal heart cycle and that can be stimulated with a pacingpulse to contract are organized in sheets and fibers that define themuscular atrial and ventricular heart chamber walls, the atrial andventricular septum, and that merge with tissues that do not contract andrelax at the base of the heart and that form valves and arterial andvenous valves, etc. The sheets and fibers forming the muscularventricles change orientation as they wrap longitudinally andtransversely around and across the atrial and ventricular walls. Thus,the cardiac tissue is anisotropic in both the intracellular andextracellular domains rather than being an isotropic medium with uniformconductivity in all directions. Moreover, the anisotropy ratio(longitudinal to transverse) is unequal in the two domains (4:1 forextracellular domain versus 10:1 for the intracellular domain).Therefore, the anisotropic cardiac tissue responds to an externallyapplied stimulus in a very interesting fashion.

Schematically illustrated responses of myocardial fibers to anodal andcathodal extracellular stimuli are illustrated in FIGS. 3A and 3B, whereL represents the longitudinal direction and T represents the transversedirection of the myocardial fibers. In FIGS. 3A and 3B, polarizedregions or fields of the myocardial fibers that are marked with a “+”are depolarized by an applied stimulus, and polarized regions or fieldsof the myocardial fibers that are marked with a “−” are hyperpolarizedby an applied stimulus. A dog-bone shaped region of depolarizationmarked “−” and extending in the transverse direction T forms directlybeneath a stimulation electrode applying cathodal stimulation to themyocardial fibers as shown in FIG. 3A. Two regions of hyperpolarizationmarked “+” also form at a distance in the longitudinal direction L awayfrom and flanking the dog-bone shaped region of depolarization marked“−” in response to the applied cathodal stimulation as also shown inFIG. 3A. The depolarization and hyperpolarization regions are reversedin polarity in response to an anodal stimulus applied through the sameelectrode, such that the central hyperpolarized region marked “+” isflanked by two regions of depolarization marked “−” as shown in FIG. 3B.

These complex hyperpolarization and depolarization patterns appearbecause the tissue anisotropy alters the flow of current in theintracellular and extracellular domains compared to that in an isotropicmedium. A way to conceptualize the two polarization fields flanking thedog-bone shaped field is to think of them as arising from “virtualsources” or “virtual electrodes”. These virtual sources/electrodes thencan be viewed as polarizing the tissue just like stimuli applied to thetissue by real electrodes at the remote polarization fields. Thepresence of these virtual sources that were initially predictedtheoretically (See Sepulveda et al., “Electric and Magnetic Fields FromTwo-Dimensional Anisotropic Bisyncytia”, Biophys J. 1987;51:557–68), hasbeen experimentally verified. See, Knisley et al., “Virtual ElectrodeEffects in Myocardial Fibers”, Biophys J. 1994; 66:719–28 and Neunlistet al., “Spatial Distribution of Cardiac Transmembrane Potentials Aroundan Extracellular Electrode: Dependence on Fiber Orientation”, Biophys J.1995;68:2310–22.

Such virtual sources can arise not only from the tissue anisotropy butalso from several other factors that disrupt the flow of theintracellular or extracellular currents, e.g., fiber bending, gradientin the extracellular electrical field, and changes in intracellular andextracellular conductance. See Sobie et al., “A Generalized ActivatingFunction for Predicting Virtual Electrodes in Cardiac Tissue”, BiophysJ. 1997;73:1410–23.

The effects of a discontinuity in a single layer of cardiac cells grownon a glass substrate has been investigated by Fast et al. in “Activationof Cardiac Tissue by Extracellular Electrical Shocks—Formation of‘Secondary Sources’ at Intercellular Clefts in Monolayers of CulturedMyocytes”, Circ. Res., 1998;82:375–385. By using fluorescent means torecord electrical activity of the cardiac cells, Fast et al. showed thatregions of polarization appear around a hole (referred to asintracellular cleft) in the monolayer in response to uniform fieldstimulus. Since these regions of polarization appear in the absence of areal electrode at the site of the hole, they can be said to be arisingfrom “virtual sources” we introduce above. Herein, we demonstrate how adiscontinuity in the intracellular domain, particularly a discontinuityin intracellular conductivity, can alter intracellular current flow andgive rise to virtual sources that can be exploited to reduce pacingthresholds.

For simplicity, consider a fiber of cardiac tissue with zero restingpotential as shown in FIG. 4A. Moreover, consider that the fiber isstimulated with anodal stimulation, for example, applied through a verysmall diameter extracellular electrode at stimulating site x=0. Thisanodal stimulation applied from a real electrode raises theextracellular potential φ_(e) and depresses the intracellular potentialφ_(l) at the site x=0 as shown in FIG. 4B. The transmembrane voltageV_(m) depicted in FIG. 4B is maximal right beneath the electrode anddecays exponentially on either side of the electrode (V_(m)=Ae^(·|x|/λ),where A is a constant that depends on intracellular and extracellularconductivities, membrane resistance, and strength of the source).

Qualitatively, this transmembrane potential V_(m) profile can beexplained as follows. As mentioned above, cardiac cells have ionchannels embedded in their membranes, and therefore have finite membraneconductance. Thus, as the current injected by the external electrodeflows along the fiber length, it flows across the cell membrane andredistributes between the extracellular and intracellular spaces asshown in FIG. 4A. The steady state current in the two domains depends onthe intracellular and extracellular conductivities (e.g. when the twoare equal the currents in the two domains are equal as well). As theintracellular and extracellular currents reach steady state,transmembrane potential V_(m) is negative (hyperpolarized) at the centerand decays to zero away from the site of the electrode (x=0).

If the intracellular space is discontinuous at a certain “spaceconstant” (λ) distance (>λ) from the electrode (at x=0) as shown by theblack mark in FIG. 5A, then the flow of intracellular current is impededas also shown in FIG. 5A. The myocardial space constant λ is a functionof intracellular resistivity (r_(e)), membrane specific resistance(r_(m)), and extracellular resistivity (r_(l)) determined by λ=√{squareroot over (r_(m)/(r_(l)+r_(e)))}. For example, the intracellular currentencounters an abrupt barrier caused by an intracellular cleft (nointracellular space) or an intracellular lesion (an induced injury tocardiac cells resulting in scar tissue), and therefore the intercellularcurrent must exit the intracellular space and reenter on the other sideof the cleft or lesion. Thus, discontinuity at the cleft or lesionimpedes the flow of intracellular current and results in changes in thetransmembrane potential V_(m) far from the electrode (x>λ) as shown inFIG. 5B.

As shown in FIG. 5C, opposite polarity regions of polarization arecreated on either side of the intracellular discontinuity or cleft orlesion at (x>λ) that would not be present in the absence of theintracellular discontinuity or cleft or lesion. The pattern oftransmembrane potential V_(m) at (x>λ) shown in FIG. 5D away from theanodal stimulus applying electrode (at x=0) can be conceptually thoughtof as arising from a pair of virtual sources that we characterize as“virtual electrodes” at the site of the intracellular discontinuity orcleft or lesion. The cathodic current exiting from the intracellularspace on one side of the cleft or lesion closest to x=0 and theaccompanied depolarization can be thought to be arising from a virtualcathodal electrode or source indicated at <−> in FIG. 5D. The anodiccurrent and the accompanying hyperpolarization on the other side of thecleft or lesion can be thought to be arising from a virtual anodalelectrode or source indicated at <+> in FIG. 4D. The virtual cathodalelectrode or source on one side of a discontinuity is referred to hereinas a “virtual cathode”, and the virtual anodal electrode or source onthe other side of the discontinuity is referred to herein as a “virtualanode”.

We found the optimal distance x to be ˜1.5 λ to ˜2.5 λ. Thetwo-dimensional spatial fields of virtual anodes and virtual cathodesoccurring on either side of an intracellular discontinuity, e.g., acleft or lesion, formed in a cardiac tissue fiber in response to ananodal stimulus, for example, applied to the cardiac tissue fiber at adistance x˜1.5–2.5 λ from the cleft or lesion are illustrated in FIGS.6A and 6B. It will be understood that the virtual anodes and virtualcathodes would appear on the opposite sides of the intracellulardiscontinuity in response to a cathodal stimulus applied to the cardiactissue fiber at the distance x˜1.5–2.5 λ from the cleft or lesion. InFIG. 6A, the elongated cleft or lesion is formed in the same directionas the longitudinal direction L of the cardiac tissue fibers. In FIG.6B, the elongated cleft or lesion is formed transverse to thelongitudinal direction L of the cardiac tissue fibers, i.e., in thetransverse direction T.

These virtual anodes and virtual cathodes flanking the cleft or lesionare conceptually similar to the hyperpolarization regions marked by “−”flanking the dog-bone shaped depolarization regions marked by “+” belowthe real cathodal electrode that are also depicted in FIGS. 6A and 6B.As described above, the regions of hyperpolarization flanking thecentral depolarized region of anisotropic cardiac tissue stimulated withcathodal stimulation depicted in FIG. 3B serve as electrotonic currentsinks. The electrotonic current sinks limit the ability of thedepolarized region directly stimulated by the real electrode to excitethe entire cardiac tissue. As noted above, the current practice is toapply sufficient pulse energy to increase the depolarized regionsufficiently to overcome the effects of the electrotonic current sinks.

By contrast, the virtual anode and virtual cathode flanking the cleft orlesion that arise when a stimulus energy is applied at the realelectrode site x=0 (FIG. 5) occur entirely due to the presence of thecleft or region and have an entirely different effect upon thedepolarization threshold.

First of all, hyperpolarization regions operating as source-sinks asdescribed above with respect to FIGS. 3A and 3B are not associated withthe virtual anode and virtual cathode bracketing the cleft or lesionillustrated in FIGS. 6A and 6B. Conceptually, this can thought to be theresult of cancellation of such effects from the closely spaced virtualanode on one side of the cleft or lesion and the virtual cathode on theother side of the cleft or lesion.

Moreover, the electrotonic interaction between the virtual anode and thevirtual cathode bracketing the lesion will be small, provided the cleftor lesion depicted in FIG. 6A is long enough (i.e., longer than spaceconstant λ). Thus, the electrotonic current sink is minimal at thelocation of the virtual cathode, and the strength of the cathodicvirtual source, and consequently the accompanied depolarization couldpotentially be stronger than the real source. As a result, the pacingthreshold would be reduced, resulting in a reduced pacing pulse energyof a pacing pulses applied at the real electrode sufficient to capturethe heart.

Similarly, the lesion may decrease the electrotonic load on thecentrally depolarized region and help reduce the pacing threshold if thelesion is made perpendicular to the fibers (i.e. along the transversedirection T) approximately at the location of the hyperpolarizationregion, as shown in FIG. 6B.

Thus, we postulated that both longitudinal and transverse lesions couldpotentially lower the pacing threshold when a pacing stimulus is appliedthrough the real electrode at x=0 in FIG. 6A or in FIG. 6B.

Experimental Methods

Experiments were performed on eight isolated guinea pig hearts weighingbetween 150 and 250 grams. Each guinea pig was anesthetized with anintraperitoneal injection of sodium pentobarbital (0.1 ml/100 g, AbbottLabs, North Chicago, Ill.), and its heart was removed via a radicalmedial thoracotomy once it failed to respond to the paw pinch reflextest. The heart was then quickly placed in a beaker containing ˜50 ml ofoxygenated 1.8 mM Ca²⁺ Tyrode (solution) maintained at ˜0° C. Thebeating heart was gently massaged to eject blood from the heart cavity,and quickly mounted on a Langendorff column. FIG. 7 shows the schematicof our experimental setup. During the entire course of the experimentthe heart was retrogradely perfused with a 1.8 mM Ca²⁺ Tyrode solutionmaintained at 36° C.–37° C. The composition of Ca²⁺ free Tyrode (in mM)was: 135 NaCl, 5.4 KCl, 1 MgCl₂, 0.33 NaH₂PO₄, 5 HEPES, 5 glucose(adjusted to pH 7.4 with NaOH). Glucose (5 mM) and Bovine Serum Albumin(BSA) (1 mg/ml) was added to the solution immediately beforeexperimentation. The heart was completely submerged in the warmed Tyrodesolution, and allowed to stabilize for ˜15–20 minutes before startingexperimentation. The solution temperature was continuously monitoredduring the stabilization period to ensure that it was within theacceptable range of 36° C.–37° C. The heart's electrical activity wasalso monitored continuously during the course of experimentation using apair of monitoring electrodes. One monitoring electrode was placed incontact with the heart wall, and the other monitoring electrode wasplaced in the bath as shown in FIG. 7. ECG type signals were derivedfrom the monitoring electrodes and used to monitor health of the heartand to determine unipolar and bipolar control pacing thresholds andpost-lesion pacing thresholds as explained below.

The unipolar and bipolar pacing electrodes used to stimulate the heartto determine pacing thresholds had surface areas of 1.2 mm² and weremade of porous platinum black, a material similar to one used onpace/sense electrodes of commercially available pacing leads. The pacingelectrodes were mounted on a manual micro-manipulator and pressedagainst the myocardium until a reliable and stable capture response topacing stimuli was observed.

The three sites on the heart where unipolar and bipolar pacingthresholds were measured comprise an anterior left ventricular siteshown in FIG. 8A, a posterior left ventricular site shown in FIG. 8B,and an interventricular septum site shown in FIG. 8C. The rightventricle was cut open along the interventricular connection to accessthe septum.

During bipolar stimulation, both the active cathodal and the returnanodal pacing electrodes were affixed at each of the three sites at aninter-electrode spacing ˜5 mm apart. During unipolar stimulation, thelead from the stimulator was removed from the return anodal pacingelectrode and connected to another return anodal electrode in the bath.The unipolar and bipolar control pacing thresholds were measured foreach site after the initial stabilization period.

A lesion was then formed in the myocardial tissue approximately midwaybetween the two pacing electrodes affixed at each site as shown. Thelesion was formed using a scalpel to make a cut ˜3–5 mm long andorthogonal to an imaginary line between the two pacing electrodesaffixed at each site. The lesion was therefore formed between the twoelectrodes at a distance l=˜2.5 mm from each electrode, resulting in aninter-electrode distance of 21=5 mm. The distance l was about ˜1.5–2.5 λ(the space constant) for the cardiac tissue. The unipolar and bipolartest pacing thresholds were then measured for each site. In someexperiments, the stability of the determined test pacing threshold wasmonitored for 5 minutes (n=7 hearts and n=21 sites) and 10 minutes(n=3hearts and n=6 sites) durations after the lesion formation.

The unipolar and bipolar, control and test, pacing thresholds weredetermined by applying a train of constant current pulses (each 0.5 msin duration with inter-pulse duration of 300 ms) using a Bloomstimulator to each site. The pulse current amplitude was graduallyincreased in increments of 0.02 mA from a sub-threshold value untilcapture of the myocardium occurred as revealed by the ECG recordings forevery pulse in the pulse train. The current amplitude at which adelivered pulse achieved capture was labeled as the pacing threshold forthat site.

For statistical comparison, a two-tailed Student's paired t-test wasused to compare the means of various data sets. The statisticalcorrelation between any two parameters was determined by calculatingPearson's correlation coefficient R, and conducting a two-tailedStudent's t-test for rejecting the null hypothesis that the slope of thebest fit line was zero, and that the parameters were not correlated.Values of P<0.05 were considered to be significant.

Experimental Results

Both unipolar and bipolar pacing threshold showed a decrease after thelesion formation (n=26 sites for both electrodes). The mean pacingthreshold using unipolar electrodes decreased by ˜50% from a controlvalue of 0.31±0.13 mA to 0.16±0.08 mA (P<0.0001). The mean pacingthreshold using bipolar electrodes also decreased by ˜50% from 0.33±0.15mA to 0.16±0.08 mA (P<0.0001). For both unipolar and bipolar electrodes,only 33% of the sites showed an increase in threshold over 5 minutes,and 20% showed an increase over 10 minutes. The worst-case increase was0.08 mA from the control value.

As described above, a train of sub-threshold amplitude pacing pulses wasapplied to the heart to explore the pacing threshold. The recorded ECGreflected sinus rhythm of the heart as shown in FIG. 9A as long as thetrain of sub-threshold pacing pulses did not capture the myocardium. Thepulse amplitude was then gradually raised in increments of 0.02 mA.Myocardial capture occurred for some pulses in the pulse train but notfor all as shown in FIG. 9B when the mean pulse amplitude was slightlybelow the threshold value. However, myocardial capture occurred forevery pulse as shown in FIG. 9C when the pulse amplitude was raised byanother 0.02 mA, and this pulse amplitude was recorded as the pacingthreshold for that site.

The reduction in pacing threshold after lesion formation is shown inFIG. 10A for unipolar stimulation. The data includes measurement from 26sites in 8 guinea pigs. The threshold was found to decrease after lesionformation for all sites except one. The mean pacing threshold decreasedfrom 0.31 mA (standard deviation SD=0.13) to 0.16 mA (SD=0.08) after thelesion formation (P<0.0001). FIG. 10B shows the same data afternormalizing the control threshold to 100%. In some experiments thereduction in pacing threshold was up to 70%.

The findings for bipolar stimulation were similar to those for unipolarstimulation. The pacing threshold decreased from 0.33 mA (SD=0.15) to0.16 mA (SD=0.08) after the lesion formation as shown in FIG. 11A(P<0.0001). The normalized data shown in FIG. 11B shows that reductionin pacing threshold for some sites was up to 80%, slightly larger thanthat observed for unipolar stimulation.

The unipolar and bipolar pacing thresholds measured 5 minutes and 10minutes after the lesion formation are shown in FIGS. 12A and 12B,respectively. After the 5 minutes wait, 13 sites (˜62%) showed adecrease in the pacing threshold, and 7 sites (˜33%) showed an increasein the pacing threshold. For unipolar stimulation, the increase in thepacing threshold was restricted to a small range of 0.08 mA for allsites except one exceptional site where the pacing threshold changed by˜0.25 mA. Possible reasons for such a large change in the pacingthreshold are discussed below. The pacing threshold increased at onlyone site after 10 minutes wait, and this increase was less than 0.02 mA.Similar data as set forth in FIG. 12B was obtained in response tobipolar stimulation, except that no unusual increase in pacing thresholdwas observed. Table 1 depicts percentage of sites with variation inpacing threshold from the control value by various fixed amounts. Note,that the variation in pacing threshold ranged from −0.1 mA to 0.08 mAafter 5 minutes wait, and the variation ranged from −0.08 mA to 0.02 mAafter 10 minutes wait.

TABLE 1 Percentage of total number of sites Pacing Threshold UnipolarBipolar Change (mA) 5 min* 10 min 5 min* 10 min 0.1 0.08 4.7% 0.06 9.6%4.7% 0.04 9.5% 14.3%  0.02 9.5% 20.0% 14.3%  20.0% −0.02 38.1%  40.0%47.6%  40.0% −0.04 9.5% 20.0% 9.5%  0.0% −0.06 4.8%  0.0% 0.0% 20.0%−0.08 0.0% 20.0% 4.8% 20.0% −0.1 9.5% n = 21 n = 5 n = 21 n = 5 *5 mindata exludes one data point with unusually large increase in pacingthreshold

The reduction in pacing threshold showed a trend of being slightlylarger for sites with higher baseline pacing threshold as shown in FIGS.13A and 13B. The reduction in pacing threshold increased with anincrease in the control pacing threshold for both unipolar electrodes asshown in FIG. 3A and bipolar electrodes as shown in FIG. 3B (r=0.36 forunipolar and r=0.44 for bipolar). Although this trend was observed forboth unipolar and bipolar stimulation, the slope was greater for bipolarstimulation (−63.4%/mA for unipolar stimulation versus −59.7%/mA forbipolar stimulation).

The pacing thresholds for the unipolar and bipolar stimulation before(control) and after the lesion formation were compared. The controlpacing threshold for the two electrodes were approximately equal(0.31±0.13 mA for unipolar versus 0.30±0.17 mA for bipolar; P<0.33) asshown in FIG. 14A. After the lesion formation the pacing threshold forthe two electrodes reduced but again remained approximately equal asshown in FIG. 14B (0.16±0.18 mA for unipolar versus 0.17±0.08 mA forbipolar; P<0.17).

These studies of guinea pig hearts support our premise thatfield-induced virtual anodes and cathode bracketing a linear lesion canreduce the pacing threshold significantly. For unipolar electrodes, themean pacing threshold was found to decrease from 0.31 mA to 0.16 mA(FIG. 11A), and for bipolar electrodes, the mean pacing threshold wasfound to be reduced from 0.33 mA to 0.16 mA (FIG. 12A). For some sites,the reduction in pacing threshold was found to be up to 70% for unipolarelectrodes (FIG. 11B) and up to 80% for bipolar electrodes (FIG. 12B).The pacing thresholds were found to be quite stable for the majority ofthe measurement sites for up to 5 to 10 minutes (FIGS. 12A and 12B andTable 1). The reduction in pacing threshold was slightly larger forhigher values of baseline pacing threshold for both unipolar and bipolarelectrodes (FIGS. 13A and 13B). And finally, for the electrode spacingused in this study, the threshold reduction for unipolar and bipolarelectrodes were found to be quite similar (FIGS. 14A and 14B). Thus, wefound that field-induced virtual sources around a linear lesion reducethe pacing threshold by ˜50% on an average. However, for some sites, areduction of up to 75–80% was observed suggesting that potentiallylarger reductions in pacing threshold are possible.

Although we observed a consistent decrease in pacing threshold for mostof the sites, this reduction showed significant variation (FIGS. 10 and11) that may be due to less than ideal control of inter-electrodedistance, lesion size and angular orientation to the lesion with respectto tissue fibers and other factors. Although attempts were made tocontrol the temperature within 36–37° C. range, it is possible that insome experiments the temperature exceeded this range for a brief period.Any such fluctuations in solution temperature will cause changes intissue excitability, and alter the action potential duration. Forexample, an increase in temperature will make the tissue hyperexcitablebecause of faster activation (See Nagatomo et al., “TemperatureDependence of Early and Late Currents in Human cardiac Wild-type andLong Q-T DeltaKPQ Na+ Channels”, Am J Physiol. 1998; 275:H2016–24) andincrease conductance of sodium channels (See Milburn et al., “TheTemperature Dependence of Conductance of the Sodium Channel:Implications for Mechanisms of Ion Permeation”, Receptors Channels1995;3:201–211) which can decrease the pacing threshold. In contrast, aslight decrease in temperature would cause a prolongation of the actionpotential.

Considering that the train of pulses used to stimulate the heart had aninter-pulse duration of 300 ms, and guinea pig action potentials aretypically 250 ms in duration, this would imply that some of the pulsescould have occurred during a relative or absolute refractory period ofthe action potential. See Watanabe et al., “Ventricular ActionPotentials, Ventricular Extracellular Potentials, and the ECG of GuineaPig”, Circ. Res. 1985;57:362–373.

Thus, higher amplitude pulses would be required to capture myocardiumwith every pulse and pacing threshold would be higher. Any othertemporal changes in the tissue condition (e.g. its metabolic state) thatresult in prolongation of action potential would result in an increasein pacing threshold by a similar mechanism.

As is clear from FIGS. 10A–10B and 11A–11B, the control pacing thresholdfor unipolar and bipolar electrodes varied from 0.1 mA to 0.7 mA.Typically, pacing threshold for normal myocardium should be in the rangeof 0.3–0.6 mA as derived from voltage and impedance values reported byHidden-Lucet et al., “Low Chronic Pacing Thresholds of Steroid-ElutingActive-Fixation Ventricular Pacemaker Leads: A Useful Alternative toPassive-Fixation Leads”, Pacing Clin Electrophysiol. 2000; 23:1798–800.Abnormally low pacing threshold for some hearts might have occurredbecause the hearts experienced transient global ischemia during theextraction procedure. During ischemia the intercellular gap junctionsare partially closed resulting in an increase in intracellularresistivity (r_(e)). Consequently the electrotonic loading effect on thetissue directly depolarized by the electrode is reduced. This wouldresult in a reduction in the critical mass, and therefore in pacingthreshold.

The percent decrease in the pacing threshold in hearts with low pacingthreshold was found to be smaller compared to the hearts with highercontrol values as shown in FIGS. 13A–13B. A plausible explanation forthis can obtained if we consider that reduction in baseline thresholdmight be correlated with severity of ischemia experienced by a heartduring the extraction procedure. An increase in intracellularresistivity (r_(e)) during ischemia implies that the myocardial spaceconstant [λ=√{square root over (r_(m)/(r_(i)+r_(e)))}; where r_(m) isthe membrane specific resistance and r_(l) is the extracellularresistivity] is decreased. Physically, this can be understood in termsof a greater resistance to the flow of intracellular current, which nowmust exit to the extracellular space over a much shorter distance fromthe stimulating electrode. As a result, intracellular and extracellularcurrents equilibrate over a shorter distance, and the space constant isreduced. In our experiments, the lesion was formed in between the twoelectrodes at a distance l=˜2.5 mm from each electrode, resulting in aninter-electrode distance 2l; which was ˜5 mm. A decrease in spaceconstant implies that the ratio of lesion-to-electrode distance to thespace constant, i.e., l/λ) increased during ischemia. As a result, thesteady state intracellular current had a greater opportunity to spreadinto surrounding myocardium before reaching the lesion. This wouldreduce the intracellular current density at the lesion. Since virtualsources arise as a result of perturbation or impediments in the flow ofintracellular current, a smaller current density implies an attenuatedstrength of virtual sources.

We found reduction in pacing threshold for unipolar and bipolarelectrodes to be quite similar. This may be the result of the fact thatdistance between the two electrodes (5 mm) was several times larger thanthe space constant (˜1 mm) of the normal myocardium. Thus, anyelectrotonic interaction between the two electrodes is expected to beminimal, i.e., current flow pattern from one electrode is unlikely to beinfluenced by the other electrode. The inter-electrode distance shouldbe of the order of space constant λ for a significant electrotonicinteraction to occur. However, if the two electrodes were to be a spaceconstant λ apart, then the distance between the lesion to any oneelectrode will be only half a space constant 0.5 λ. This would beinsufficient distance for intracellular current density to reach asteady state maximal value. Thus, to maximize the strength of thevirtual sources, the two electrodes should at least be two spaceconstants away (2 λ). However, this guarantees that electrotonicinteraction between the two electrodes will be small or negligible, andtherefore the unipolar and bipolar electrodes should yield identicalresults as observed in this study.

Pacing Lead Embodiments

The present invention can be embodied in epicardial and endocardialpacing leads of the types known in the prior art. A first embodiment ofa pacing lead that both forms a discontinuity and provides pacingstimulation is depicted in FIGS. 15–19. The endocardial pacing lead 10comprising a lead body 26 extending between a proximal connectorassembly 20 and a distal electrode head 50. A stylet 30 is also depictedin FIG. 15 having an elongated stylet wire 32 extending from a styletknob 34 and inserted down the lumen of the lead body 26.

Lead body 26 is formed of a length of outer insulating sheath 12 havingproximal and distal ends and a sheath lumen, the sheath 12 operating asan electrical insulator formed of a biocompatible silicone rubber orpolyurethane compound substantially inert to body fluids. A multi-filar,coiled wire conductor 18 having proximal and distal ends and a coillumen formed therein is loosely received within the sheath lumen ofsheath 12.

The proximal connector assembly 20 comprises a connector ring 22 and aconnector pin 24 that are electrically connected to separately insulatedanodic and cathodic wires of the multi-filar coiled wire conductor 18.Sealing ring sets 28 and 28′ are compressed and serve to seal the leadbody lumen and the gap between the connector pin 24 and connector ring22 from ingress of body fluids upon insertion of connector assembly 20into a mating bore of an implantable pulse generator connector block ina manner well known in the art.

Sheath 12 and coiled wire conductor 18 extend between the connectorassembly 20 and the electrode head 50 shown in cross-section in FIGS. 18and 19. The electrode head 50 can include a plurality of soft, plianttines 52 (shown in FIG. 15 but omitted from FIGS. 16–19 for convenienceof illustration) that provide passive fixation of the electrode head 40disposing the electrode head distal surface 54 against the endocardiumin a manner well known in the art. Anodic and cathodic pacing electrodes56 and 58 and a cutting element or blade 60 extend distally from theelectrode head distal surface 54. The anodic and cathodic pacingelectrodes 56 and 58 are electrically connected through the separatelyinsulated anodic and cathodic wires of the multi-filar coiled wireconductor 18 to the connector ring 22 and connector pin 24,respectively.

The electrode head distal surface 54 is preferably non-conductive andsupports the anodic and cathodic pacing electrodes 56 and 58 on eitherside of a centrally disposed cutting element or blade 60. The anodic andcathodic pacing electrodes 56 and 58 are spaced apart by aninter-electrode distance 2l; which may be on the order or 3.0 mm to 5.0mm apart. The anodic and cathodic pacing electrodes 56 and 58 can beformed of any of the known pacing materials and have an electrodesurface area of about 1–2 mm² or 1.2 mm² for example. The pacingelectrodes 56 and 58 can be formed on the surface of the electrode headdistal surface 54 or can project distally from the surface of theelectrode head distal surface 54. The pacing electrodes 56 and 58 arecoupled through conductors 66 and 68, respectively, to wires of thecoiled wire conductor 18.

The cutting element or blade 60 is adapted to be retracted into thedistal electrode head 50 during implantation as shown in FIG. 19 andejected distally of the electrode head distal surface 54 at theimplantation site to cut through the endocardium and form thediscontinuity as shown in FIG. 18. The cutting element or blade 60 ispreferably formed of a non-conductive ceramic or a similar material thatcan have a highly sharpened cutting edge 62. The cutting edge 62preferably extends about 2 mm further distally when ejected as shown inFIG. 18 than the anodic and cathodic pacing electrodes 56 and 58.

The electrode head 50 of lead 10 is advanced via a percutaneous accessinto a vein with the cutting element or blade 60 retracted into a slit64 and cylindrical chamber 70 of the electrode head 50 as shown in FIG.19. A first stylet 30 having stylet wire 32 length that extends to thedistal electrode head 40 can be used to stiffen and steer the lead body26 during implantation.

A second stylet 30 having a stylet wire 32 length that extends throughthe length of the distal electrode head 40 can then be used to distallyadvance the cutting element or blade 60 from the distal electrode head40 to form a lesion or cleft in endocardial tissue. The proximal end ofthe cutting blade 60 is mounted to a mounting block 76 embedded within asealing block 80 having a plurality of sealing rings 78 bearing tightlyagainst the surface of the cylindrical chamber 70. A proximal recess 74of the movable mounting block 76 can be engaged by the distal surface ofstylet wire 32. The stylet wire 32 is advanced distally within the leadbody lumen to engage the recess 74 to move the assembly of the sealingblock 80, the mounting block 76, and the cutting blade 60 distally fromthe retracted position of FIG. 19 to the extended position of FIG. 18.

The movement of the cutting blade 60 is done with sufficient force topenetrate through the cardiac tissue layers. The cutting edge 62 isshaped like the edge of a razor blade. The electrode head is firmlyadvanced against the cardiac tissue in initial implantation, and thetines hold the advanced position.

The present invention may also be employed in an epicardial pacing leadwhere it may not be necessarily necessary to move cutting element orblade. The electrode head 150 of an exemplary epicardial pacing lead 100is depicted in part in FIGS. 20 and 21. The lead body 126, coiled wireconductor 118, and proximal connector assembly (not depicted) can takethe forms employed in the endocardial lead 10, except that a styletlumen and stylet are not necessary. The electrode head 150 is providedwith a mesh 112 on the electrode head surface 154 adapted to encouragetissue ingrowth and that can be sutured through or adhered to theepicardium employing a medical adhesive following procedures known inthe prior art.

Again, the anodic and cathodic pacing electrodes 156 and 158 are spacedapart by an inter-electrode distance 2l; which may be on the order or3.0 mm to 5.0 mm apart. The anodic and cathodic pacing electrodes 156and 158 can be formed of any of the known pacing materials and have anelectrode surface area of about 1–2 mm² or 1.2 mm² for example. Thepacing electrodes 156 and 158 can be formed on the surface of theelectrode head distal surface 154 or can project distally from thesurface of the electrode head distal surface 154. The pacing electrodes156 and 158 are coupled through wires of the coiled wire conductor 118to the connector pin and ring of the proximal connector assembly.

The movement of the cutting blade 160 is done with sufficient force topenetrate through the cardiac tissue layers. In this embodiment, thecutting blade 160 is shaped like an arrow that has a point 164 thatfirst penetrates the tissue layer before the arrow edge 162 extendingaway from the point 164 cuts through the cardiac tissue.

An electrode head 250 of a further exemplary epicardial pacing lead 200that supports a solid, non-conductive, screw 260 to form a discontinuityin cardiac tissue between the anodic and cathodic pacing electrodes 256and 258 is depicted in FIGS. 22 and 23. The lead body 226, coiled wireconductor 218, and proximal connector assembly (not depicted) can takethe forms employed in the endocardial lead 10, except that a styletlumen and stylet are not necessary.

Again, the anodic and cathodic pacing electrodes 256 and 258 are spacedapart by an inter-electrode distance 2l, which may be on the order or3.0 mm to 5.0 mm apart. The anodic and cathodic pacing electrodes 256and 258 can be formed of any of the known pacing materials and have anelectrode surface area of about 1–2 mm² or 1.2 mm² for example. Thepacing electrodes 256 and 258 can be formed on the surface of theelectrode head distal surface 254 or can project distally from thesurface of the electrode head distal surface 254. The pacing electrodes256 and 258 are coupled through wires of the coiled wire conductor 218to the connector pin and ring of the proximal connector assembly.

The electrode head 250 is provided with a mesh 212 adapted to encouragetissue ingrowth and that can be sutured through or adhered to theepicardium employing a medical adhesive following procedures known inthe prior art. Or, fixation may be possible by the threads 262 of solidhelix or screw 260 that are intended to be screwed into the myocardiumemploying a fixation tool, e.g., the tool disclosed in commonly assignedU.S. Pat. No. 6,010,526. The electrode head 250 is grasped by the tooland pressed against the epicardium as the tool and electrode head 250are rotated so that sharpened tip 264 penetrates the epicardium, and thethreads 262 screw into the myocardium and draw the epicardium againstthe pacing electrodes 256 and 258.

Although an epicardial lead is depicted in FIGS. 22 and 23, it will beunderstood that the solid helix or screw 260 can be substituted for thecutting blade 60 of the endocardial lead of FIGS. 15–19.

In each such embodiment, the virtual anode and virtual cathode arecreated in the cardiac tissue on the sides of the blade 60 or 160 or thescrew 260 that are closest to the cathodic and anodic pacing electrodesas depicted in the figures.

It will be understood that the above-described embodiments areparticularly useful for bipolar pacing as shown or for unipolar pacing,wherein the indifferent anodal pacing electrode 56, 156, 256 is notpresent or is not employed by the IPG.

CONCLUSION

All patents and publications identified herein are incorporated hereinby reference in their entireties.

While particular embodiments of the invention have been disclosed hereinin detail, this has been done for the purposes of illustration only, andis not intended to limit the scope of the invention as defined in theclaims that follow. It is to be understood that various substitutions,alterations, or modifications can be made to the disclosed embodimentwithout departing from the spirit and scope of the claims. The abovedescribed implementations are simply those presently preferred orcontemplated by the inventors, and are not to be taken as limiting thepresent invention to the disclosed embodiments. It is therefore to beunderstood, that within the scope of the appended claims, the inventionmay be practiced otherwise than as specifically described withoutactually departing from the spirit and scope of the present invention.

1. A method of applying pacing pulses to cardiac tissue having a spaceconstant to effect depolarization of a cardiac tissue mass of a heartchamber comprising; forming a discontinuity in the cardiac tissue of theheart chamber; disposing a pacing electrode supported by an electrodehead at a distance exceeding the space constant of the cardiac tissuefrom the discontinuity in the cardiac tissue; and applying a stimulus tothe cardiac tissue though the pacing electrode to form a depolarizationregion at the pacing electrode insufficient to propagate adepolarization wave front through the cardiac tissue mass of the heartchamber but sufficient to induce a transmembrane potential change at thetissue adjacent to the discontinuity that results in a propagateddepolarization wave front to effect a contraction of the heart chamber,wherein the step of forming a discontinuity comprises forming anintracellular cleft by applying a cutting blade to the cardiac tissue tocut the cardiac tissue, wherein the cutting blade is supported to extendfrom the electrode head.
 2. The method of claim 1, wherein the cardiactissue comprises cardiac cells that are elongated in length and arrangedin anisotropic cardiac fibers and the step of forming a discontinuitycomprises forming an elongated discontinuity in the cardiac tissueextending in alignment with or transverse to the lengths of theelongated cardiac cells.
 3. The method of claim 1, wherein the distanceexceeding the space constant is about 1.5–2.5 times the space constant.4. A method of applying pacing pulses to cardiac tissue having a spaceconstant to effect depolarization of a cardiac tissue mass of a heartchamber comprising; forming a discontinuity in the cardiac tissue of theheart chamber; disposing a pacing electrode supported by an electrodehead at a distance exceeding the space constant of the cardiac tissuefrom the discontinuity in the cardiac tissue; and applying a stimulus tothe cardiac tissue though the pacing electrode to form a depolarizationregion at the pacing electrode insufficient to propagate adepolarization wave front through the cardiac tissue mass of the heartchamber but sufficient to induce a transmembrane potential change at thetissue adjacent to the discontinuity that results in a propagateddepolarization wave front to effect a contraction of the heart chamber,wherein the step of forming a discontinuity comprises forming anintracellular cleft by applying an electrically insulated cutting bladeto the cardiac tissue to cut the cardiac tissue and retaining theelectrically insulated cutting blade in the intracellular cleft, whereinthe electrically insulated cutting blade is supported to extend from theelectrode head.
 5. A method of applying pacing pulses to cardiac tissuehaving a space constant to effect depolarization of a cardiac tissuemass of a heart chamber comprising; forming a discontinuity in thecardiac tissue of the heart chamber; disposing a pacing electrodesupported by an electrode head at a distance exceeding the spaceconstant of the cardiac tissue from the discontinuity in the cardiactissue; and applying a stimulus to the cardiac tissue though the pacingelectrode to form a depolarization region at the pacing electrodeinsufficient to propagate a depolarization wave front through thecardiac tissue mass of the heart chamber but sufficient to induce atransmembrane potential change at the tissue adjacent to thediscontinuity that results in a propagated depolarization wave front toeffect a contraction of the heart chamber, wherein the cardiac tissuecomprises cardiac cells that are elongated in length and arranged inanisotropic cardiac fibers and the step of forming a discontinuitycomprises forming an elongated discontinuity in the cardiac tissueextending in alignment with or transverse to the lengths of theelongated cardiac cells, wherein the step of forming a discontinuitycomprises forming an intracellular cleft by applying an electricallyinsulated cutting blade to the cardiac tissue to cut the cardiac tissueand retaining the electrically insulated cutting blade in theintracellular cleft, wherein the electrically insulated cutting blade issupported to extend from the electrode head.
 6. A cardiac pacing lead ofthe type comprising an elongated cardiac lead body extending between aproximal electrode connector and a distal electrode head for applyingpacing pulses to cardiac tissue having a space constant to effectdepolarization of a cardiac tissue mass of a heart chamber comprising:discontinuity forming means supported on the electrode head for forminga discontinuity in the cardiac tissue of the heart chamber; and at leastone pacing electrode supported by the electrode head and disposed at adistance exceeding the space constant of the cardiac tissue from thediscontinuity forming means through which a pacing pulse can bedelivered at an energy insufficient to cause the directly stimulatedtissue adjacent to the pacing electrode to propagate a depolarizationwave front through the cardiac tissue mass of the head chamber butsufficient to induce a transmembrane potential change at the tissueadjacent to the discontinuity that results in a propagated wave front toeffect a contraction of the heart chamber, wherein the distanceexceeding the space constant is about 1.5–2.5 times the space constant.7. The cardiac pacing lead of claim 6, wherein the discontinuity formingmeans comprises means for forming a lesion in the cardiac tissue.
 8. Thecardiac pacing lead of claim 6, wherein the discontinuity forming meanscomprises a non-conductive cutting blade having a point supported toextend from the electrode head to be inserted into the cardiac tissue tocut the cardiac tissue.
 9. The cardiac pacing lead of claim 6, whereinthe discontinuity forming means comprises means fix forming anintracellular cleft in the cardiac tissue.
 10. The cardiac pacing leadof claim 6, wherein the discontinuity forming means comprises a solid,non-conductive, fixation screw supported to extend from the electrodehead to be screwed into the cardiac tissue.
 11. The cardiac pacing leadof claim 6, wherein the cardiac tissue comprises cardiac cells that areelongated in length and arranged in anisotropic cardiac fibers and thediscontinuity forming means comprises means for forming an elongateddiscontinuity extending in alignment with or transverse to the lengthsof the elongated cardiac cells.
 12. The cardiac pacing lead of claim 11,wherein the discontinuity forming means comprises means for forming anelongated intracellular cleft in the cardiac tissue.
 13. The cardiacpacing lead of claim 11, wherein the discontinuity forming meanscomprises a non-conductive cutting blade having a point supported toextend from the electrode head to be inserted into the cardiac tissue tocut the cardiac tissue.
 14. The cardiac pacing lead of claim 11, whereinthe discontinuity forming means comprises a non-conductive cutting bladesupported to extend from the electrode head to be inserted into thecardiac tissue to cut the cardiac tissue.
 15. The cardiac pacing lead ofclaim 11, wherein the discontinuity forming means comprises means forforming a lesion in the cardiac tissue.
 16. A cardiac pacing lead of thetype comprising an elongated cardiac lead body extending between aproximal electrode connector and a distal electrode head for applyingpacing pulses to cardiac tissue having a space constant to effectdepolarization of a cardiac tissue mass of a heart chamber comprising:discontinuity forming means supported on the electrode head for forminga discontinuity in the cardiac tissue of the heart chamber; and at leastone pacing electrode supported by the electrode head and disposed at adistance exceeding the space constant of the cardiac tissue from thediscontinuity forming means through which a pacing pulse can bedelivered at an energy insufficient to cause the directly stimulatedtissue adjacent to the pacing electrode to propagate a depolarizationwave front through the cardiac tissue mass of the heart chamber butsufficient to induce a transmembrane potential change at the tissueadjacent to the discontinuity that results in a propagated wave front toeffect a contraction of the heart chamber, wherein the discontinuityforming means comprises a non-conductive cutting blade supported toextend firm the electrode head to be inserted into the cardiac tissue tocut the cardiac tissue.